Radiation-based imaging systems, mainly X-ray systems, are routinely used for diagnostic examination purposes prior to interventional procedures e.g. in cardiology, radiology and surgery. Such systems generally consist of a C-arm on which are mounted a radiation source and a radiation detector, i.e. an X-ray tube assembly and an X-ray detector for example, a high-voltage generator for generating the tube voltage, an imaging system including monitor, a control device and a patient positioning table. Systems having two C-arms that are movable in separate planes are also known.
These days flat-panel detectors are typically used as X-ray radiation detectors. Normally such flat-panel detectors are indirect-converting detectors in which the incident radiation that has penetrated the examination object is not converted directly into electrical signals, but is initially converted into light, which is then converted into electrical signals. Detectors of said type therefore comprise a scintillator layer consisting of elongated needles made from a scintillator material forming said scintillator layer. Said scintillator layer is applied on a substrate. Photooptically coupled to the scintillator layer is an active readout pixel array consisting of a multiplicity of photodiodes arranged in a matrix shape, each photodiode being associated with one pixel. Each pixel also includes a switching element, typically in the form of a transistor, the individual pixels obviously being provided also with corresponding drive and readout electronics. This pixel and readout array can be implemented on the basis of CMOS or related technologies, while detectors having active photodiode and readout arrays composed of polycrystalline silicon are also known.
The scintillator layer can consist for example of CsJ, which creates the scintillator layer in the form of tightly packed needles that have been grown on the substrate. Other scintillator materials, such as Gd2S2O, CuJ, CsF, BaF2, CeF3, BGO for example, are also known and structured in similar fashion.
During operation the X-rays, after having penetrated the object, strike the scintillator layer and, depending on the hardness of the radiation, i.e., the radiation intensity, are absorbed in different planes in the scintillator layer and converted into light, which is to say that the incident X-ray quanta are converted into light quanta. Said light quanta are guided by way of the scintillator needles to the optically coupled pixel array located there under, where they are incident on the individual photodiodes. There, the light quanta are converted into electrical signals, which are then read out.
The basic structure and the basic principle of operation of such a radiation detector are sufficiently well-known.
The quantum efficiency of a scintillator varies depending on radiation quality, e.g., between roughly 50%-80%, dependent on radiation quality, in the case of a scintillator made of CsJ having a layer thickness of, e.g., 600 μm. As a result the spatial frequency-dependent detective quantum efficiency DQE(f) (DQE=Detective Quantum Efficiency) is limited at the upper end and is even significantly below this for typical pixel sizes of, e.g., 150-200 μm and for the applications of relevant spatial frequencies of 1-2 lp/mm. Above the K-edge it basically holds that the absorption decreases with increasing radiation hardness, and as a consequence thereof so too does the DQE(f).
Harder radiation occurs very frequently in interventional cardiology applications, for example. The treatment of corpulent patients may be cited as an example, as also may the recording of oblique projections, a penetration of the patient of up to 40 cm and more being necessary in both cases, which is to say that the X-ray quanta travel a correspondingly long way, namely up to 40 cm and more, through the patient. Due to the decreasing DQE(f) it is not always possible in this case to deliver images of acceptable quality, while the X-ray dose requiring to be applied is also considerable.